There is a great need for flexible, multi-scale, configurable sensor/effector systems that can be cut to specific sizes and shapes and adapted “on the fly” to adjust to specific temporal and spatial recording scales in the body. A driving motivation for such technology is the rapidly growing awareness that brain, cardiac and other biological recordings must be made multi-scale, spanning from individual and multiple unit (cell) activities to large-scale field potentials, depending upon the application. These sensor systems must also be capable of both recording and modulating tissue function, as part of new diagnostic and therapeutic devices for neurological and other diseases, brain and other tissue injury, and acquired conditions.
Taking an example from the brain-computer interface technologies, the current state of the art for flexible subdural grid electrodes for localizing seizures in the human brain typically utilize ˜4 mm platinum-iridium or stainless steel contacts spatially separated by 10 mm.
Such electrodes are available from Ad-Tech Medical Instrument Corporation (http://www.adtechmedical.com). As illustrated in FIG. 1, a long-term monitoring (LTM) subdural grid 10 of this type containing 16 contacts 20 has a separation of 10 mm between each contact 20, and 1 or 2 “tails” 30 that contain contacts 40 for output of electrical signals that correspond to each contact 20 on the subdural grid (each contact 20 on the subdural grid has an individual wire attached to it, which is connected to a contact 40 in one of the “tails.”) Similar electrodes are illustrated in U.S. Pat. No. 4,735,208, entitled “Subdural strip electrodes for determining epileptogenic foci.” However, these electrodes are not effective for detecting all signals of interest in the brain, for example, due to their large size and large spacing between contacts.
The choice of the particular sizing and spacing of the electrodes is partly based upon clinical tradition and partly based on technological limitations of the design, such as requiring that each contact has a wire dedicated to conducting its signals to the recording apparatus. Studies on neocortical neuron density suggest that there are approximately 12 million neurons contained within the square centimeter of neocortex sampled by each one of these electrodes (See Pakkenberg B & Gundersen H J. Neocortical neuron number in humans: effect of sex and age. J Comp Neurol (1997) 384: pp. 312-320). It seems very unlikely that this would be a sufficient spatial sampling to capture even a small amount of the information available from this type of recording. The exact resolution for sensing and modulating activity in brain, peripheral nerve, spinal cord or other tissues in the body depends upon the particular application (e.g. brain-computer interface, functional electrical stimulation, alleviation of pain, etc.). A single electrode system that has the capability to resolve a broad range of these activities, tunable to a particular task, would be highly desirable and useful. In addition, being able to configure dimensions of the recording surface, through a broad range of configurations (e.g. to interface with a particular gyms, dorsal root entry zone, peripheral or cranial nerve bundle) would be highly desirable, and contribute to great economy in power use, computational burden and minimize disruption of normal tissues.
In addition, the actual tissue contact region of the electrode system needs to be “changeable,” as different applications may require recording from either the surface of tissues, from contacts that penetrate tissues to be close to particular types of cells, nuclei, nerve bundles or specific tissues, or perhaps from a combination or adjustable array of contact types. The proposed system is designed specifically with this type of flexibility in mind, allowing the “business end” of the system, the portion of the system that actually contacts biological tissues, to be adaptable and changeable in many different combinations.
To achieve the desired range of spatial sampling for signals of interest along with the desired area of coverage, it is clear that the number of sensor/effector contacts must be on the order of thousands, not tens or even hundreds, and that the spatial resolution of these sensor/effector contacts must be “scalable,” (e.g. their effective size and spacing be adjustable without having to physically move or alter them). As an example, since each contact of existing brain/subdural electrode systems is either individually wired and assembled (e.g. Ad-Tech systems) or fabricated such that individual wires output signals from each contact (e.g. Utah array), it is clear that a more integrated design that incorporates multiplexing control techniques is needed to minimize the number of leads required and to make production feasible and cost effective. This also provides a safety advantage over current intracranial electrode systems, as there is evidence that the number of “tails” or leads extruding from the body can be related directly to morbidity in the case of subdural grids for brain recording, for example.
In the case of subdural electrodes for monitoring brain activity, other electrode designs have attempted to overcome the problem of spatial undersampling by making the electrodes smaller and more closely spaced. For example, the Utah Electrode Array 50 shown in FIG. 2 has an array of contacts spaced 0.4 mm apart. The Utah Electrode Array is described by Nordhausen C T, Maynard E M & Normann R A in “Single unit recording capabilities of a 100 microelectrode array,” Brain Res. (1996), Vol. 726, pp. 129-140. While this provides a more desirable density of electrodes, the overall area of cortex that is sampled by the electrode array 50 is only 4 mm×4 mm, due to the small array size of 10×10 contacts. This amount of spatial coverage is insufficient for most clinical applications. Extending this electrode design to a larger array size is difficult because each electrode must be individually wired and because the array is made from inflexible silicon that does not conform to the shape of the tissues.
Several improvements have been suggested to fix the first problem of wiring complexity. Two such examples are illustrated in FIGS. 3 and 4. The example of FIG. 3 is described by Patterson W, Yoon-Kyu Song, Bull C, Ozden I, Deangellis A, Lay C, McKay J, Nurmikko A, Donoghue J & Connors B. in “A microelectrode/microelectronic hybrid device for brain implantable neuroprosthesis applications,” IEEE Transactions on Biomedical Engineering (2004), Vol. 51, pp. 1845-1853, while the example of FIG. 4 is described by Aziz J, Genov R, Bardakjian B, Derchansky M & Carlen P. in “256-channel integrated neural interface and spatio-temporal signal processor, Circuits and Systems, 2006. ISCAS 2006. Proceedings. 2006 IEEE International Symposium on (2006), p. 4. In these electrode circuit designs, each electrode is connected to its own amplifier cell 60 (inset, FIG. 4). As shown in FIG. 4, each cell 60 includes a programmable high pass filter 61 and low pass filter 62, preamp 63, final amp 64, a sample-and-hold circuit 65, and an analog memory 66. In this way, each electrode can have its own dedicated amplifier and programmable filter bank. The outputs of all of the amplifier cells 60 in a given column are multiplexed together using an array of analog switches 70, and rows are multiplexed using an analog multiplexer 80 to allow all of the electrode outputs to be reduced to a single time-division multiplexed output line 85. This technique greatly reduces the number of wires that must exit the electrode array. However, the inflexible silicon substrate that these circuits are fabricated on still limits their use to sampling a small area of brain tissue where the surface can be approximately flat.
An ideal sensor/effector array would be flexible and stretchable to allow it to conform to the round and contoured surface (and within sulci and other recesses) of the brain or other biological tissues. Some attempts have been made to fabricate implantable electrodes using flexible printed circuit technology. For example, FIG. 5 illustrates the implantable electrodes 90 described in U.S. Pat. No. 6,024,702. However, this technique only allows for passive circuit elements, and so the problem of wiring complexity still remains.
The present invention addresses these and other needs in the art.